JP3376005B2 - Magnetic resonance imaging - Google Patents

Description

Description: BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a magnetic resonance imaging apparatus. In particular, the present invention relates to a technique for obtaining magnetization or spin-lattice relaxation time in a thermal equilibrium state in a subject. 2. Description of the Related Art Magnetic resonance imaging is a technique in which a group of nuclei having a unique magnetic moment is placed in a uniform static magnetic field.
This is a phenomenon in which the energy of a high-frequency magnetic field rotating at a specific frequency is resonantly absorbed, and this technique is used to visualize chemical and physical phenomena of a substance. Among these physical phenomena, spin-nucleon relaxation and spin-spin relaxation are important phenomena, and these phenomena are used in many diagnoses. This is a normal part and diseased part, spin describe the process of this relaxation - nucleon relaxation times T ₁ and spin -
Spin relaxation time T ₂ are different. Apart from this, it is necessary to measure T ₁ and T ₂ also from the viewpoint of quantification. This is because, normally, under the conditions of a pulse sequence for imaging, the magnetic resonance signal is affected by T1 and T _2, for the quantification because shall correction of T ₁ and T ₂ is there. When measuring T ₁ , IR (Inver
sion Recovery) method is used. FIG. 3 shows a pulse sequence of the IR method. This method uses a pre-pulse 18
In this method, the magnetization in the thermal equilibrium state is reversed in the -Z direction by a 0-degree pulse, and a 90-degree pulse, which is an observation pulse, is applied after a time _t1i to perform observation. By changing the time t _1i between the pre-pulse and the observation pulse, a magnetization recovery process as shown in FIG. 4 is obtained, and M ₀ and T ₁ are obtained by curve fitting.
It is a method of seeking. However, in this method, it is necessary to wait until the magnetization M ₀ in the thermal equilibrium state is restored after the application of the 90-degree pulse.
The interval T _D of degree pulse had to be set to 5 times the T _1. Therefore, the observation time Tobs is expressed by the following equation, and there is a problem that it takes a long time. In addition,
In Equation 1, n is the number of t _1i . [0005] Therefore, in order to solve the above problem, FI
The R (Fast Inversion Recovery) method was developed. In this method, in order to shorten the observation time, T _{D is set} to T ₁ of 2 to 3 times, and the recovery process of magnetization can be obtained by changing the time t 1 _i of the pre-pulse and the observation pulse as in the IR method. is there. Fitting is performed on the recovery process thus obtained, and M ₀ and T ₁ are obtained. If T _D = 2T ₁ , Tobs is given by the following equation. [0007] However, this FIR method also requires an observation time corresponding to T _D. In contrast, in the SR (Saturation Recovery) method with 90 degrees prepulse, it is possible to set the T _D almost 0, the observation time is as shown in the following equation. The pulse sequence of this SR method is shown in FIG.
Shown in [0009] As is apparent from the above equations 2 and 3,
In the SR method, the observation time can be shortened as compared with other methods. However, as shown in FIG. 6, the change in magnetization is 0 to M ₀ , which is almost １ ／ compared to that of the FIR method.
Become. For this reason, the SR method has a problem that accuracy is deteriorated. On the other hand, there are several methods for measuring T ₂ as described below. That is, one of the T ₂ measurement methods
One is a method in which measurement is performed while changing the echo time, and the obtained several signals are obtained by processing such as curve fitting. The pulse sequence used at this time is shown in FIG.
It is shown in FIG. The change of the echo signal can be described by T ₂ , and the echo signal sequence is expressed by using the echo time TE.
It can be represented by (-TE / T2). However,
In this method, data must be collected every time the echo time is changed, so that there is a problem that an observation time is required. Further, when the TE is lengthened, the influence of the spin diffusion becomes more liable to occur, which deviates from the curve of exp (-TE / T2), and there is a problem that the measurement accuracy is lowered. As a solution to this problem, Call and Pu
There is known a method called a CP method in which a 90-degree pulse is connected to a 180-degree pulse after the 90-degree pulse, such as 90-180-180-. FIG. 15 shows this pulse sequence, and FIG.
Shown in In this pulse sequence, the spin has a 9
Since it is affected only between 0 degrees and 180 degrees or between 180 degrees and 180 degrees, the influence of diffusion can be reduced by shortening these intervals. However, this method has a problem that it is susceptible to the imperfection of the 180-degree pulse. Meiboom and Gill are 90 degrees X ′ to 180 degrees y ′ to 18 that change the phases of 90 degrees and 180 degrees by 90 degrees.
CPMG (Call-P
urcell-MeibooM-Gill) method was devised. FIG. 17 shows this pulse sequence. In this method, odd-numbered echoes are affected by imperfections in the high-frequency magnetic field, while even-numbered echoes are not. Therefore, performs curve fitting using only even-numbered echo signal intensity, by obtaining a T _2, effects of imperfections of the radio frequency magnetic field is not subjected. This is shown in FIG. However, there is a problem that the odd-numbered echo signal cannot be used for curve fitting, and data of the odd-numbered echo is wasted. On the other hand, the measurement of the T ₂ distribution can be performed by a method combining this CPMG method with phase encoding and frequency encoding. However, the odd-numbered echo signals Like the T ₂ measurement method in this imaging method there is a problem that wasted. In addition, the quantification has been mentioned earlier, and the Gibbs ringing phenomenon also poses a problem in quantification. The Gibbs ringing phenomenon is a phenomenon in which a voxel has a finite size, so that a signal is mixed into another voxel. This Gibbs ringing phenomenon can be explained by band limitation in the frequency space. The density distribution of the substance is shown in FIG.
Consider the case where When this is Fourier-transformed, it becomes as shown in FIG. However, since the voxel size is finite, the band is limited on the k-space as shown in FIG. 3C, and the data actually collected on the k-space is data as shown in FIG. . FIG. 4B is Fourier transformed, and FIG. 4C is Fourier transformed, and FIG. 4C is Fourier transformed, as shown in FIG. 4F. (G) in which FIG. As a result, Gibbs ringing occurs and a signal is mixed into other voxels. This effect becomes very problematic when the high signal region and the low signal region are close to each other. For example, head imaging of creatine phosphate or adenosine triphosphate, a 31P metabolite, can be mentioned.
In muscle, the content of creatine phosphate and adenosine triphosphate is about 7 to 8 times that of the brain. For this reason, the measurement accuracy in the brain deteriorates due to the mixing of muscle signals into the brain,
It was a problem. As described above, the conventional FIR method can measure with high accuracy, but has a disadvantage that the observation time is long. On the other hand, in the SR method, although the observation time can be shortened, there is a disadvantage that accuracy is deteriorated. On the other hand, according to the conventional T2 measurement method, the collected odd-numbered data cannot be used, and this is wasteful.
There was a problem that the measurement accuracy was reduced. Further, in the conventional magnetic resonance imaging apparatus, when a high signal area and a low signal area are in the vicinity, there is a problem that an image cannot be accurately formed in a low signal area due to mixing of signals due to Gibbs ringing. According to the present invention, in order to solve the above-mentioned problems, a high-frequency magnetic field and a gradient magnetic field are applied to a subject placed in a uniform static magnetic field. In a magnetic resonance imaging apparatus for obtaining a magnetic resonance image by collecting generated magnetic resonance signals, a first pulse in which a high-frequency magnetic field pulse in which a first pulse is a 180-degree pulse and a second pulse is a 90-degree pulse is applied. An application unit, a second application unit that applies a high-frequency magnetic field pulse in which the first pulse and the second pulse are composed of 90-degree pulses, and a pulse generated from the subject based on the first or the second application unit. Means for collecting magnetic resonance signals and obtaining first or second reconstructed data; and at least one of magnetization or spin-lattice relaxation time in a thermal equilibrium state based on the combined data obtained by the means. Means for determining one of them, wherein the first applying means applies a high-frequency magnetic field pulse when a time interval between the first pulse and the second pulse is equal to or less than a predetermined time, and the second applying means The magnetic resonance imaging apparatus is characterized in that a high frequency magnetic field pulse is applied when a time interval between the first pulse and the second pulse is equal to or longer than the predetermined time. Further, according to the present invention, a high frequency magnetic field and a gradient magnetic field are applied to a subject placed in a uniform static magnetic field, thereby collecting a magnetic resonance signal generated from the subject to obtain a magnetic resonance image. In the magnetic resonance imaging apparatus, a first high-frequency magnetic field pulse having a pulse angle of 90 degrees is applied 90.
High-frequency magnetic field pulse applying means, and a 180-degree high-frequency magnetic field for applying a plurality of second high-frequency magnetic field pulses having a phase difference of 90 degrees with respect to the first high-frequency magnetic field pulse and a pulse angle of approximately 180 degrees A pulse applying unit, a unit for collecting a magnetic resonance signal generated from the subject corresponding to each of the second high-frequency magnetic field pulses, and an odd-numbered application of the 180-degree high-frequency magnetic field pulse applying unit. A magnetic resonance imaging apparatus comprises a means for calculating a spin-spin relaxation time based on each of a corresponding magnetic resonance signal and a magnetic resonance signal corresponding to an even-numbered application. Further, according to the present invention, a high-frequency magnetic field and a gradient magnetic field are applied to a subject placed in a uniform static magnetic field, so that magnetic resonance signals generated from the subject are collected to obtain a magnetic resonance image. In the magnetic resonance imaging apparatus, of the plurality of voxels constituting the magnetic resonance image, a means for calculating a mixed amount of an image signal from one voxel to another voxel, and based on the mixed amount obtained by this means. Means for obtaining a corrected image from the one voxel image signal constitutes a magnetic resonance imaging apparatus. According to the first aspect of the invention, the portion where the time between the pre-pulse and the observation pulse is short is obtained by the FIR method, and the magnetization inverted in the -Z direction can be observed.
The portion where the time between the pre-pulse and the observation pulse is long can be obtained by the SR method, and the magnetization in the Z direction can be observed. That is, the change in the magnetization to be recovered can be increased to approximately twice M ₀ as in the FIR method.
It can be made equivalent to the IR method. Also, the FIR method that requires a waiting time T _D after the observation pulse is used only in the portion where the time between the pre-pulse and the observation pulse is short, and the SR method that can make T _D almost zero is used for the long portion. Measurement can be performed in a short observation time. That is, with the above configuration, it is possible to accurately obtain M ₀ and T ₁ . Embodiments of the present invention will be described below with reference to the drawings. FIG. 2 is a block diagram showing the configuration of the magnetic resonance imaging apparatus according to one embodiment of the present invention. In the drawing, a static magnetic field magnet 1 and a gradient coil 2 and a shim coil 4 provided inside the static magnetic field magnet 1 apply a uniform static magnetic field to a subject (not shown) and linearly move in three directions x, y, and z orthogonal to each other in the same direction. A gradient magnetic field having a gradient magnetic field distribution is applied. The gradient coil 2 is driven by a gradient coil power supply 5, and the shim coil 4 is driven by a shim coil power supply 6. Gradient coil 2
The probe 3 provided inside the probe applies a high-frequency magnetic field to the subject by receiving a high-frequency signal from the transmission unit 7 and receives a magnetic resonance signal from the subject. The probe 3 may be used for both transmission and reception, or may be provided separately for transmission and reception. The magnetic resonance signal received by the probe 3 is detected by the receiving unit 8 and then transferred to the data collecting unit 9 where it is A / D converted and sent to the computer system 10 for data processing. The above-described gradient coil power supply 5, shim coil power supply 6, reception unit 8 and data collection unit 9 are all controlled by a pulse sequence control unit 12, and the pulse sequence control unit 12 is controlled by a computer system 10. The computer system 10 is controlled by a command from the console 11. From the data collection unit 9 to the computer system 1
The magnetic resonance signal input to 0 is subjected to a Fourier transform or the like, and image data of a density distribution of a desired nucleus in the subject is reconstructed based on the Fourier transform. This image data is sent to the image display 13 and displayed as an image. Next, a method for obtaining M ₀ and T ₁ will be described. As the pulse sequence, a pulse sequence of the FIR method shown in FIG. 7 and a pulse sequence of the SR method shown in FIG. 5 are used. In the FIR method, the signal magnitude Mobs follows Equation 4, and the relationship between Mobs and t1i is as shown in FIG. (Equation 4) In the SR method, the magnitude of the signal in accordance with the following formula, Mobs and t _1i
Is as shown in FIG. (Equation 5) In the present invention, in order to increase the change in magnetization, a short t _1i
For, the data is collected by the pulse sequence of the FIR method, and for long t _1i , the pulse sequence of the SR method is used. FIG. 1 shows the recovery process of the magnetization at this time.
It is. This gives the observation time as
It can be shorter than the FIR method. (Equation 6) Here, the model formula is set as follows. (Equation 7) Using this model formula, fitting is performed by the nonlinear least squares method. In this method, the amount of change in the magnitude of the magnetization can be approximately doubled as in the FIR method, so that the same accuracy as in the FIR method can be obtained. Next, a case where the flip angle of the high frequency magnetic field is not 90 degrees or 180 degrees will be described. In this case, equation 8 is used for the FIR method, and equation 9 is used for the SR method.
Obey. [Equation 8] (Equation 9) As in Equation 8 or Equation 9, h relating to the flip angle is added to the parameter. In this case, Equation 10 is used instead of the model equation shown in Equation 7 above. [Mathematical formula-see original document] M ₀ and T ₁ are obtained by fitting according to this model formula. In order to image the T ₁ distribution by this method, the pulse sequence shown in FIG. 9 or FIG. 11 is used for the FIR method, and the pulse sequence shown in FIG. 10 or FIG. A sequence may be used. Next, a method of measuring T ₂ will be described. FIG. 17 is a diagram showing a pulse sequence of the CPMG method. This embodiment uses this pulse sequence.
When the pulse deviates from 180 degrees, the spin behaves as follows in the CPMG method. Where 180
The deviation from the degree is assumed to be α, and a (180-α) degree pulse will be described. First, the spin falls on the y 'axis of the rotating coordinate system by the 90-degree pulse (FIG. 19A). Next, the gap spreads as shown in FIG. 19B due to the non-uniformity of the magnetic field until the interval τ between the 90 ° pulse and the (180−α) pulse. Here, the (180-α) degree pulse applied in the y ′ direction causes the spin to be inverted where it floats from the x′y ′ plane (FIG. 3C). Thereafter, the spin moves in the y'-axis direction and gathers at a position floating from the y'-axis as shown in FIG. This is the first echo. At the time of τ until the next (180-α) degree pulse, it moves as shown in FIG.
-Α) The degree pulse is applied (FIG. 9 (f)). And
The spin moves again in the y′-axis direction, and after τ time, the spins gather on the y′-axis (FIG. 9G). This is the second echo. In this way, the odd-numbered echoes gather at a position floating from the y 'axis, while the even-numbered echoes gather on y'. Therefore, the echo signal sequence is as shown in FIG. 18, and the even-numbered echo signal Meven (TE) is expressed by the following equation. [Equation 11] On the other hand, the odd-numbered echoes are gathered at a position slightly floating from the y'-axis, but the odd-numbered echo signals Modd
(TE) is also attenuated at T ₂ and can be expressed by the following equation with k as a proportional constant. (Equation 12) In order to use both these even-numbered echoes and odd-numbered echoes, the data is arranged as shown in FIG. In this data array, the model formula of Expression 11 is used in the region of FIG. 5A, and the model formula of Expression 12 is used in the region of FIG. Using this model formula, M
A non-linear least squares method is performed using ₀ , k, and T ₂ as parameters. Thus, it is possible to use both the even-numbered echo and the odd-numbered echoes, thus improving the T ₂ measurement accuracy. The method of measuring T ₂ has been described above. On the other hand, T
_In order to obtain the _two distributions, a pulse sequence shown in FIG. 20 or FIG. 21 is used. With this pulse sequence,
An echo signal train as shown in FIG. 18 is obtained for each pixel. These signal strings are converted into data strings as shown in FIG. After that, curve fitting is performed for each pixel using the model equations shown in Equations 11 and 12. With this method, the T ₂ distribution can be determined. Finally, a correction method in metabolite imaging will be described with reference to FIG. First, a case where a high signal region and a low signal region are known in advance in metabolite imaging will be described. For example, in ³¹ P metabolite imaging, creatine phosphate and adenosine triphosphate correspond to this. The difference between these metabolites in muscle and brain is about 7 to 8 times, and it has been previously known that the effect of Gibbs ringing is great. In this case, the position of the muscle is confirmed by the 1 H image, and the amount of the signal mixed into the brain may be calculated. First, the distribution in the voxel of the metabolite image is determined based on the 1 H image (step 1). However, the voxel for which the distribution is to be obtained may be a voxel in which the brain and the muscle coexist as shown in FIG. 23A, or a voxel in which the muscle exists only in a part of the voxel as shown in FIG. For these voxels, the distribution is obtained by dividing the voxel into M × M × M. FIG. 24 shows the correspondence between voxel coordinates and coordinates after subdivision. However, only the one-dimensional direction is shown. Next, in step 2, the mixing amount is calculated. here,
The coordinates of the high signal voxel are (p, q, r), and the voxel of the mixing destination is (n, 1, m). The density distribution in the high-signal voxel required for calculating the mixing amount can be expressed by Expression 13. Here, the voxel signal can be considered as a muscle signal based on the density difference between the muscle and the brain. Therefore, the muscle signal of each pixel is obtained by dividing the voxel signal value by the number of pixels including the muscle. [Mathematical formula-see original document] [Mathematical formula-see original document] The following equation is used to calculate the amount of contamination. [Mathematical formula-see original document] In Expression 15, M is an even number here. Next, correction is performed based on the obtained mixing amount to obtain a corrected image (step 3). As for the correction method, the low signal voxel is obtained by subtracting the mixing amount from the signal value. The high-signal voxel is obtained by adding the amount of mixing into the low-signal voxel to the signal value. This is the end of the correction. Next, another embodiment will be described. First, a voxel in which the image signal ratio of adjacent voxels in the metabolite image is equal to or more than a is searched for, and a high-signal voxel is found. a is set in advance. Next, a high signal area on the 1H image corresponding to the high signal voxel is confirmed. In the correction method described above with reference to FIG. 22, this high signal region corresponds to a muscle signal. That is, the processing of FIG. 22 performed on the previous muscle signal is performed on this high signal area. As described above, a corrected image can be obtained. In the method described above, the distribution of metabolite images in high-signal voxels was determined as being uniform. However, it is also possible to determine only the portions corresponding to these voxels by separately forming high-resolution metabolite images. . With this method, the distribution within the voxel can be obtained, and the amount of contamination can be calculated and corrected by the method of FIG. As described above, according to the present invention, it is possible to accurately obtain T ₁ in a short observation time.

BRIEF DESCRIPTION OF THE DRAWINGS FIG. 1 is a diagram showing a process of recovering magnetization when the present method is used. FIG. 2 is a block diagram of a magnetic resonance imaging apparatus. FIG. 3 is a diagram showing a pulse sequence of the IR method. FIG. 4 is a diagram showing a recovery process of magnetization in the IR method. FIG. 5 is a diagram showing a pulse sequence of the SR method. FIG. 6 is a diagram showing a magnetization recovery process in the SR method. FIG. 7 is a diagram showing a pulse sequence of the IR method. FIG. 8 is a diagram showing a recovery process of magnetization in the FIR method. FIG. 9 is a diagram showing a pulse sequence for obtaining a T ₁ distribution according to an embodiment of the present invention. FIG. 10 is a diagram showing a pulse sequence for obtaining a T ₁ distribution according to an embodiment of the present invention. FIG. 11 is a diagram showing a pulse sequence for obtaining a T ₁ distribution according to an embodiment of the present invention. FIG. 12 is a diagram showing a pulse sequence for obtaining a T ₁ distribution according to an embodiment of the present invention. FIG. 13 is a diagram for describing an example in which an echo signal sequence collected by the CPMG method is divided into even-numbered echoes and odd-numbered echoes and arranged in data. FIG. 14 is a view showing a sequence of a high-frequency magnetic field pulse for spin echo. FIG. 15 is a diagram showing a pulse sequence for performing the CP method. FIG. 16 is a diagram showing a state in which an echo signal sequence by the CP method is curve-fitted. FIG. 17 is a diagram showing a pulse sequence for performing the CPMG method. FIG. 18 is a diagram showing a state in which an echo signal sequence by the CPMG method is curve-fitted. FIG. 19 is a view for explaining spin behavior in the CPMG method. FIG. 20 is a diagram showing a pulse sequence for obtaining a T ₂ distribution according to an embodiment of the present invention. FIG. 21 is a diagram showing a pulse sequence for obtaining a T ₂ distribution according to an embodiment of the present invention. FIG. 22 is a flowchart showing a method for correcting the effect of Gibbs ringing. FIG. 23 is a diagram showing a voxel distribution in a metabolite image. FIG. 24 is a diagram showing the coordinates of voxels and subdivided voxels. FIG. 25 is a diagram for explaining the Gibbs ringing phenomenon. [Description of Signs] 1 ... Static magnetic field magnet 2 ... Gradient coil 3 ... Shim coil 4 ... Probe 5 ... Gradient coil 6 ... Shim coil power supply 7 ... Transmission unit 8 ... Reception unit 9 ... Data collection unit 10 ... Computer system 11 ... Console 12 ... Pulse sequence control unit 13: Image display

Claims (1)

(57) Claims 1. A high-frequency magnetic field and a gradient magnetic field are applied to a subject placed in a uniform static magnetic field to collect magnetic resonance signals generated from the subject. A magnetic resonance imaging apparatus for obtaining a magnetic resonance image, wherein a first pulse is a 180-degree pulse and a second pulse is a high-frequency magnetic field pulse composed of a 90-degree pulse.
An application means for applying a high-frequency magnetic field pulse in which the first pulse and the second pulse are composed of 90-degree pulses; and a second pulse generated from the subject based on the first or second application means. Means for collecting magnetic resonance signals to be obtained and obtaining first or second reconstructed data; means for obtaining at least one of magnetization in a thermal equilibrium state or spin-lattice relaxation time based on the combined data obtained by the means. Wherein the first application unit applies a high-frequency magnetic field pulse when a time interval between the first pulse and the second pulse is equal to or less than a predetermined time, and the second application unit includes the first pulse A high-frequency magnetic field pulse is applied when a time interval between the second pulse and the second pulse is equal to or longer than the predetermined time.